A compact drum kit comprising a body having a hollow interior defined by a top, a bottom, and three or more sides connecting the top and bottom, a sound hole, a plurality of component mounting shafts mounted in the hollow interior of the body and protruding through a plurality of openings in the top, a plurality of leg shafts mounted in the hollow interior of the body and protruding through a plurality of openings in the bottom, one or more percussion modules that attach to one or more of the plurality of component mounting shafts, wherein each percussion module produces a sound in response to being struck or scraped, and a kick pedal including a beater positioned to strike a first side of the three or more sides and produce a sound.

A drum kit, drum set, trap set, or just drums is a collection of drums and other percussion instruments set up to be played/struck by a single player. The traditional drum kit consists of a mix of drums (classified as classically as membranophones, Hornbostel-Sachs high-level classification 2) and idiophones (Hornbostel-Sachs high-level classification 1, most significantly cymbals but also including the woodblock and cowbell for example). More recently kits have also included electronic instruments (Hornbostel-Sachs classification 53), with both hybrid and entirely electronic kits now in common use.
A standard modern kit (for a right-handed player), as used in popular music and taught in many music schools, contains: A snare drum, mounted on a stand, placed between the player's knees and played with drum sticks (which may include rutes or brushes); a bass drum, played by a pedal operated by the right foot; and one or more cymbals, played with the sticks.
All of these are classed as non-pitched percussion, allowing for the music to be scored using percussion notation, for which a loose semi-standardized form exists for the drum kit. If some or all of them are replaced by electronic drums, the scoring and most often positioning remains the same, allowing a standard teaching approach. The drum kit is usually played seated on a drum stool or throne. The drum kit differs in general from those traditional instruments that produce melody or chords/pitch per se: even though drums are often placed musically alongside others that do, such as the piano or stringed instruments.
Many drummers extend their kits from this basic pattern, adding more drums, more cymbals, and many other instruments including pitched percussion. In some styles of music particular extensions are normal, for example double bass drums in heavy metal music. On the other extreme but more rarely, some performers omit elements from even the basic setup, also dependent on the style of music and individual preferences.


A method of non-invasively measuring venous oxygen saturation, comprising: applying a pressure transducer at a first site on a body, applying a drive signal to the external pressure transducer at a predetermined frequency, to cause a series of pulsations of a predetermined magnitude in the venous blood volume in the vicinity of said first site, applying an oximeter device at a second site on the body, measuring output signals received from said oximeter device, said output signals containing a component representative of the modulation of venous blood volume due to said pulsations, deriving a measure of venous oxygen saturation from the frequency response of said output signals.

The monitoring of venous blood oxygen saturation (SvO2) is called venous oximetry. The method is often used in Intensive Care Units (ICU) to monitor the patient's overall oxygen supply and consumption. Current invasive methods have resulted in its under-utilization although SvO2 is unquestionably a valuable assessment tool in the evaluation of oxygenation.
All venous oximetry techniques can be categorised into two areas, methods that are invasive and those that are non-invasive. A discussion of the various known invasive and non-invasive techniques follows.
Oxygen saturation can be measured invasively by employing a variation of the standard Swan-Ganz Pulmonary Artery Catheter (PAC) in which two fiberoptic bundles were inserted in the PAC. The modified PAC used the principle of reflection spectrophotometry to make quantitative measurement of oxygen transport. Due to its invasive nature and the cost of the modified PAC, the method is not employed extensively for venous oximetry. U.S. Pat. No. 5,673,694 is representative of this background art.
The majority of non-invasive, continuous peripheral venous oximetry techniques are based on Near InfraRed Spectroscopy (NIRS) or a combination of NIRS and various exercise protocols such as over-systolic venous occlusion. NIRS is hindered from supplanting the current invasive, continuous method utilizing SvO2 or central venous catheter mainly due to the difficulty in determining certain critical parameters without which calibration for venous oximetry would not be possible. Prior art approaches based upon NIPS are disclosed in U.S. Pat. Nos. 6,015,969 and 5,661,302.
Pulse oximetry is one of the main applications of photoplethysmography (PPG), and is widely used for the measurement of arterial oxygen saturation, SpO2. The PPG waveform contains two components, one which is attributable to the pulsatile component in the vessels, i.e. the arterial pulse, is caused by the heartbeat and gives a rapidly alternating signal (AC component), and the other is due to the blood volume and its change in the skin and gives a steady signal that only changes slowly (DC component).
Two wavelengths of light are used in Pulse oximetry, one in the red band (660 nm) and one in the infrared band (940 nm). Since at 660 nm reduced hemoglobin absorbs more light than oxyhemoglobin and at 940 nm, oxyhemoglobin absorbs more light than its reduced form, pulse oximetry relates this differential measurement to the arterial oxygen saturation.
In pulse oximetry, light is first being transmitted through the tissues and the intensity of the transmitted light is then measured by a photo detector on the other side. The pulse oximeter first determines the AC component of the absorbance at each wavelength and then divides it by the corresponding DC component to obtain ratio that is independent of the incident light intensity. The ratio of ratios is then constructed as:
  R  =                    AC        660            /              DC        660                            AC        940            /              DC        940
The pulse oximeter is then calibrated by measuring the ratio of ratios and simultaneously sampling arterial blood for in vitro saturation measurements.
Whilst the use of these techniques is effective for the measurement of arterial blood oxygen saturation, it relies upon the presence of pulsations of the arterial blood which is generated by the heart. No such measurable pulsations are present in venous blood.
Venous Occlusion Plethymography (VOP) is the measurement of changes in tissue volume in response to temporary obstruction of venous return. It is used clinically to measure certain physiological conditions of blood vessels such as venous capacitance. VOP relies on the principle that occlusion of venous return causes slight swelling of distal portion of the tissue under test due to continued arterial inflow. The step response of venous blood volume over time during VOP can be used to measure arterial blood flow, venous outflow and venous compliance.
M. Nitzal et al (Journal of Biomedical Optics 5(2), 155-162, April 2000) employed the principle of VOP to the measurement of SvO2, by applying pressure to the forearm sufficient to completely occlude venous flow, but leave arterial flow unaffected. Light absorption at 2 wavelengths is compared before and after occlusion. However, the approach does not appear to yield separate determination of venous and arterial oxygen saturation.
PCT publication WO99/62399 and U.S. Pat. No. 5,638,816 relate to methods of venous oximetry where a cyclical active pulse is applied via an external cuff. However, the level of modulation (10% of the DC signal) is large, and will require a cuff-sensor spacing so close that the optical coupling will be affected, or if further away, pulsations will be at a level which will cause perturbations in the arterial system, and hence lead to inaccuracies in venous oxygen saturation measurements.
As indicated above, prior art techniques for measuring venous oxygen saturation by non-invasive means do not yield the requisite accuracy. The aim of the invention is to achieve an improved measure of venous oxygen saturation.
The principles of arterial pulse oximetry are well known (see above). The crucial element of the method that enables specific calibration of the oxygen carrying hemoglobin depends upon the presence of blood volume pulsations in the arterial system. These pulsations are of course naturally present throughout the circulation system. If one could induce pulsations in the venous system and properly isolate them from those of the arterial system, a similar calibration method could be employed to measure venous oxygen saturation.
According to the invention, there is provided a method of non-invasively measuring venous oxygen saturation, comprising
applying a pressure transducer at a first site on a body,
applying a drive signal to the external pressure transducer at a predetermined frequency, to cause a series of pulsations of a predetermined magnitude in the venous blood volume in the vicinity of said first site,
applying an oximeter device at a second site on the body,
measuring output signals received from said oximeter device, said output signals containing a component representative of the modulation of venous blood volume due to said pulsations,
deriving a measure of venous oxygen saturation from the frequency response of said output signals.
The term oximeter device is intended to encompass any device which uses light of different frequencies to determine tissue oxygen content. It may encompass both transmission and reflection mode devices.
The relationship between the distance between the first and second sites on the one hand, and the magnitude of the pulsations on the other hand, may be arranged such that a multiplicative term in the frequency spectrum of the measured signals, indicative of a disturbance to the arterial system, is minimised.
Preferably the frequency of the drive signal is chosen such that the pulsations are distinguishable from the heart rate.
The relationship between the frequency ωm of the pulsations caused by said drive signal and the heart rate ωHR can be chosen to comply with the following conditions:ωm≠n(ωHR±ΔωHR)andωm±(ωHR±ΔωHR)≠n(ωHR±ΔωHR)
where n>1
The frequency of the drive signal can be determined iteratively on the basis of a real time measurement of the heart rate.
Ideally the magnitude of the pulsations is controlled such that the venous blood system is modulated without disturbing the arterial system. The magnitude of the pulsations can be controlled such as to cause a variation of less than 1% in the DC level of the received signal, or more preferably a variation of approximately 0.1% in the DC level of the received signal.
In an alternative embodiment the oximeter is placed on a digit and a further pressure transducer is placed away from a distal end of a limb, and arranged to occlude the supply of blood to and from said limb, and the measurement is performed during the period of occlusion.
Advantageously, the optimum magnitude of the pulsations can be determined by progressively increasing the pulsation magnitude and observing a frequency response of the output signals; the appearance of a multiplicative term in the frequency spectrum being indicative of the maximum permissible magnitude having been reached. Upon reaching the point where a multiplicative term appears, the pulsation magnitude can be subsequently reduced to a point where the multiplicative term becomes insignificant.
The method may further include a calibration step, during which the spacing between the first and second sites is varied while the pulsation magnitude is held constant, in order to derive an optimum spacing. The calibration step may include a further step of varying the pulsation magnitude while the spacing of the first and second sites is held constant.
In a further embodiment the method may include the measurement of arterial oxygen saturation derived from the frequency response of the output signals, such that the difference in levels of the arterial and venous oxygen saturation being representative of tissue oxygen consumption.
The invention also provides an apparatus for non-invasively measuring venous oxygen saturation comprising
a pressure transducer for applying a series of pulsations to a first site on a body,
a pulse oximeter,
control means for controlling the frequency and/or magnitude of said pulsations, such that the venous blood volume is modulated,
signal processing means for extracting a value for venous oxygen saturation from signals received from said oximeter, said signals containing a component representative of a modulation of venous blood volume due to said pulsations.
Advantageously, the control means operates to control the relationship between the frequency ωm of the pulsations caused by said drive signal and the heart rate ωHR according to the following conditions:ωm≠n(ωHR±ΔωHR)andωm±(ωHR±ΔωHR)≠n(ωHR±ΔωHR)
where>1
Preferably the pressure transducer comprises an inflatable digit cuff supplied with air from an air pump; the pressure transducer and the pulse oximeter may be formed as an integral device.
With reference to FIG. 1, the principle of the invention will now be described. In its simplest form, the invention involves some means for inducing changes in venous blood volume and a corresponding means for measuring the changes induced. The signals extracted are processed to yield at least a separate value for venous oxygen saturation, and where necessary, a value for arterial oxygen saturation.
The generalized theory underlying the invention will now be explained. Extending the lowest order conventional description of arterial pulse oximetry can make a zeroth order theoretical description of venous pulse oximetry. The Beer-Lambert law, which couples physical path length and effective absorbance into a single definition of optical density, is commonly used in arterial pulse oximetry to assign physical significance to changes in the optical path length. According to this model, we can write the received intensity due to a particular illuminating wavelength, λ, in terms of the proportion of arterial hemoglobin that is chemically combined with oxygen, S,I(t,λ)=I0(λ)exp{−[SεHbO2(λ)+(1−S)εHb(λ)]z(t)−μstaticd},  (1)
where εHbO2(λ), εHb(λ) are the millimolar extinction coefficients of oxygenated and de-oxygenated hemoglobin respectively, z(t) is a function of both the dynamic physical path length through arterial blood and the total hemoglobin concentration, and μstaticd is the optical density of the non-pulsatile tissue and other anatomical components.
By distinguishing optical paths through venous zv(t) and arterial za(t)blood we may generalize the model equation (1) toI(t,λ)=I0(λ)exp{−μaza(t)−μvzv(t)−μstaticd},   (2)
where we have made the substitutionsμa(λ)=└SaεHbO2(λ)+(1−Sa)εHb(λ)┘μv(λ)=[SvεHbO2(λ)+(1−Sv)εHb(λ)]
We now explore small changes in the received intensity resulting from small changes in the optical paths (resulting from the presence of low amplitude venous and arterial modulations) and consider the resultant changes (AC) normalized by the quasi-static (DC) intensity, namely
                                          Δ                                                                                I                                            (                                  t                  ,                  λ                                )                                                          I                                      (                              t                ,                λ                            )                                      ≅                                            -                              μ                a                                                  Δ                                                                                z                a                                                          (                t                )                                              -                                    μ              v                                    Δ                                                                                z                  v                                                                  (                  t                  )                                            .                                                          (        3        )
The quantities expressed in equation (3) can be separated by electronic (or other signal processing methods) since the induced venous modulations are of known origin. One method of separation is to induce a frequency modulation of the venous system in a band that is distinct from the arterial pulsations. Once isolation of the arterial and venous dynamics is achieved the process of calibration can be applied. Inversion of the classical Beer-Lambert model of pulse oximetry is usually achieved by generating two instances of equation (3) at two different wavelengths. These equations are then solved for a quantity known as the ratio of ratios, R, which is defined as the ratio of the total extinction of blood at the two wavelengths used. Assuming each term in equation (3) has been isolated, we may form two "ratio of ratios":
                                          R            a                    =                                                    [                                                                            I                                                                    (                                                  t                          ,                                                      λ                            2                                                                          )                                                                                Δ                                                                                I                                                                    (                                                  t                          ,                                                      λ                            1                                                                          )                                                                                I                                                                    (                                                  t                          ,                                                      λ                            1                                                                          )                                                                                Δ                                                                                I                                                                    (                                                  t                          ,                                                      λ                            2                                                                          )                                                                                ]                            a                        =                                                            μ                  a                                                                  (                                      λ                    1                                    )                                                                              μ                  a                                                                  (                                      λ                    2                                    )                                                                                R            v                    =                                                    [                                                                            I                                                                    (                                                  t                          ,                                                      λ                            2                                                                          )                                                                                Δ                                                                                I                                                                    (                                                  t                          ,                                                      λ                            1                                                                          )                                                                                I                                                                    (                                                  t                          ,                                                      λ                            1                                                                          )                                                                                Δ                                                                                I                                                                    (                                                  t                          ,                                                      λ                            2                                                                          )                                                                                ]                            v                        =                                                                                μ                    v                                                                          (                                          λ                      1                                        )                                                                                μ                    v                                                                          (                                          λ                      2                                        )                                                              .                                                          (        4        )
Knowledge of the extinction coefficients of oxygenated and deoxygenated hemoglobin at the two wavelengths can then be used to estimate Sa,Sv from the calculated R using the formulae
                                          S            a                    =                                                                      R                  a                                                                                ɛ                    Hb                                                                          (                                          λ                      2                                        )                                                              -                                                ɛ                  Hb                                                                  (                                      λ                    1                                    )                                                                                R                  a                                                                  [                                                                                ɛ                        Hb                                                                                (                                                  λ                          2                                                )                                                              -                                                                  ɛ                                                  HbO                          2                                                                                (                                                  λ                          2                                                )                                                                              ]                                            -                              [                                                                            ɛ                      Hb                                                                                (                                              λ                        1                                            )                                                        -                                                            ɛ                                              HbO                        2                                                                                (                                              λ                        1                                            )                                                                      ]                                                                                S            v                    =                                                                      R                  v                                                                                ɛ                    Hb                                                                          (                                          λ                      2                                        )                                                              -                                                ɛ                  Hb                                                                  (                                      λ                    1                                    )                                                                                R                  v                                                                  [                                                                                ɛ                        Hb                                                                                (                                                  λ                          2                                                )                                                              -                                                                  ɛ                                                  HbO                          2                                                                                (                                                  λ                          2                                                )                                                                              ]                                            -                              [                                                                            ɛ                      Hb                                                                                (                                              λ                        1                                            )                                                        -                                                            ɛ                                              HbO                        2                                                                                (                                              λ                        1                                            )                                                                      ]                                                                        (        5        )
The invention will now be illustrated by way of example by description of a specific embodiment, involving the modulation of venous blood volume in the index finger to thereby inject a pulsatile signal into the venous blood. The methods of achieving the modulation, recording of the signal and extraction of the modulated signal will be outlined below.
A digit cuff measuring 90 mm long and 16 mm wide was obtained from Hokanson®. Modulation of the venous blood volume within the index finger is achieved by continuously inflating and deflating the digit cuff which is wrapped around the base of the index finger. The high modulating frequency (six to seven Hertz) can be achieved mainly because the volume of air needed to inflate the digit cuff in order to cause a partial occlusion of the digit and hence a significant fractional blood volume change is small. In the same way, the time to deflate the cuff is short as the volume of air within the cuff is small.
A micro pressure air pump from Sensidyne® was obtained as an air source for the digit cuff. It is able to maintain an air flow of 6.1 LPM and a minimum pressure of 4.9 psig. These specifications are suitable for the application of inflating the digit cuff to a suitable pressure in order to cause a significant fractional blood volume change and cause a pulsatile signal that is comparable to the arterial pulsation.
A solenoid operated pinch valve was obtained from BioChem Valve® in order to modulate the digit cuff. The three-way pinch valve has one normally closed valve and one normally open valve. When the solenoid is energised, the configuration changes over. One valve is used to control the tube leading from the micro air pump to the digit cuff and the other is used to control the outlet from the digit cuff. During inflation, the tube leading from the micro air pump is opened in order to allow air to enter the digit cuff and the tube that is used as an air outlet from the digit cuff is closed. During deflation, the tube leading from the air pump into the digit cuff is 'pinched' and therefore closed and at the same time the tube leading from the digit cuff which is used as an air outlet is opened to allow air within the digit cuff to escape. By controlling the three-way pinch valve, modulation of the digit cuff at a certain frequency can be achieved.
The new method of non-invasive venous oximetry works in conjunction with a standard pulse oximeter finger probe attached to a PPG system. The digit cuff is wrapped around the base of the index finger and the finger probe is also mounted to record the modulated venous blood volume signal. The probe to cuff spacing is selected to yield a suitable level of modulation in the received signal, as will be discussed in more detail later in this specification. As the finger probe used comes equipped with two light sources, red and infrared, two different signals of the modulated venous blood volume are recorded. These signals are in turn used to formulate the ratio of ratios which is related to the oxygen saturation of the venous blood (SvO2).
The recorded signal will consist of two signals of different frequencies. One signal is the PPG signal that is related to the arterial pulsation which frequency is the subject's heart rate. The second signal is the modulated venous blood volume signal, that may be modulated at a frequency set away from the arterial pulsation so as to aid extraction of it through filtering techniques.
FIG. 2 illustrates the operation of this embodiment in schematic form. A discussion of the results obtained with this embodiment now follows.
The diagram in FIG. 3 shows the PPG signals (AC components) recorded without any venous blood volume modulation. The signal sources used to obtain the top waveform and bottom waveform were IR and Red respectively.
The diagram in FIG. 4 shows the waveforms that were recorded when the venous blood volume was modulated at a frequency higher than the PPG signal and at a pressure that was below systole.
By performing a Fast Fourier Transform (FFT) on the modulated waveforms, the power spectrum density of the two wavelengths can be obtained. FIG. 5 shows the power spectrum density at the two wavelengths.
It can be seen from the power spectrums that the modulation of venous blood was at seven Hertz and at a pressure lower than the systolic pressure. The modulated venous blood volume can be easily extracted through band-pass filtering method and the ratio of ratios can be formulated in the same way as described earlier in the background discussion of the pulse oximeter. This can then be calibrated in the same manner with measurement of the oxygen saturation of blood samples by co-oximeter drawn from the pulmonary artery. In this way, a non-invasive mixed venous blood oximetry can be achieved.
A preliminary calibration of uncalibrated SaO2 to uncalibrated SvO2 is shown in FIG. 6. The graph clearly shows a correlation between arterial oxygen saturation and venous oxygen saturation.
As mentioned above, it is important that the signal due to modulation of the venous blood can be separated from the PPG signal. One way of achieving this is to ensure that the frequency with which the pulsations are applied is chosen such that is distinct from the PPG signal.
In order to optimise the modulation frequency, both the modulation fundamental frequency and the multiplicative term if any should be set away from the second harmonic of the PPG signal. The frequency of the multiplicative term not only depends on the modulation frequency but also on the PPG fundamental frequency, that is, the heart rate. Since the normal heart rate at rest can vary, an optimisation phase can be introduced just before any actual measurement during operation. The purpose of the optimisation phase is to first gauge the heart rate of the subject and then position the modulation frequency in such a way that both the 1st and 2nd harmonics of the PPG signal are not obscured by the modulation fundamental frequency or the multiplicative term.
An optimisation phase may be needed since the variability of the heart rate and the perceived modulation depth at the measurement site depends on the physiological profile of the subject.
The upper limit of the modulation frequency is a function of the frequency response of the vascular system. Therefore too high a modulation frequency will cause the modulating signal to fail to register completely.
During this optimisation phase, the modulation depth can be adjusted according to the magnitude of the multiplicative term, as will be described in more detail later. In this way both the modulation frequency and modulation depth could be optimised before any actual measurement during operation. The optimisation phase can also be made adaptable throughout the operation, so that real time changes in heart rate could be compensated for by adapting the modulation frequency to the changes.
There are two conditions that need to be satisfied during modulation frequency optimnisation. In order to illustrate these two conditions, it helps to use a mathematical model each to represent the PPG signal and the modulating signal.
The PPG signal can be represented by the equation:
                                          f            HR                                          (            t            )                          =                              ∑                          n              =              1                                                                    f              n                                                  sin                                            (                                                      n                                                                                ω                      HR                                                            t                                    +                                      ϕ                    HR                                                  )                                                                        (        1        )
Where ƒn is the coefficient and ωHR is the PPG fundamental frequency and n represent the harmonics (n=1 is the fundamental frequency), and φHR is the phase involved.
The modulating signal can be represented by the equation:ƒm(t)=g0 sin(ωmt+φm)  (2)
Where g0 is the coefficient ωm is the modulation frequency and φm is the phase in the model.
To illustrate the variability of the heart rate the equation (1) can be rewritten as:
                                          f            HR                                          (            t            )                          =                              ∑                          n              =              1                                                                    f              n                                                  sin                                            (                                                                            n                                                                    (                                                                              ω                            HR                                                    ±                                                      Δω                            HR                                                                          )                                                                                t                                    +                                      ϕ                    HR                                                  )                                                                        (        3        )
The two conditions which need to be met for optimisation of the measurement are:ωm≠n(ωHR±ΔωHR)  (4)andωm±(ωHR±ΔωHR)≠n(ωHR±ΔωHR)  (5)
where n>1
In this way, a table can be constructed to identify the forbidden frequencies. Assuming a heart rate is 70 and the variability is ±10
Forbidden frequency bandsωm for conditionωm forat (4)condition at (5)2 Hz to 2.7 Hz1 Hz to 4 Hz3 Hz to 4.05 Hz2 Hz to 2.7 Hz
Therefore ωm should be at least greater than 4 Hz but less than the filter upper cut-off frequency for signal detection (typically 3 Hz to 6 Hz). Most commercial SpO2 devices use low cut-off frequencies. Thus ωm must be set in a narrow range of optimised band.
The graphs shown in FIGS. 7a to 7e are the power spectrums of the PPG signal with venous modulation at various frequencies. A relatively high modulation depth of 160 mmHg was chosen such that the multiplicative term could be significantly registered.
In the first graph shown in FIG. 7a, the first harmonic (n=2) of the PPG signal overlapped with the multiplicative term and the modulation fundamental overlapped with the 2nd harmonic. Thus the modulation frequency of 4 Hz failed both the conditions. This would pose a problem when it comes to spectrally separating the two signals. Furthermore, in optimising the modulation depth, the magnitude of the multiplicative term is important in determining the degree of coupling between the two signals, the overlapping of the multiplicative terms with PPG harmonics would cause erroneous measurement of the magnitude.
In the second graph (FIG. 7b), at 4.5 Hz, the multiplicative term started to move away from the first harmonic but the modulation frequency remained dangerously close to the 2nd harmonic.
The third graph illustrates the situation when the heart rate was increased. Although the modulation frequency has increased to 5 Hz, the multiplicative term and modulation fundamental still overlap with the 1st and 2nd harmonics respectively. This demonstrates that the optimisation depends on the heart rate variability as well.
The fourth graph (FIG. 7d) illustrates the situation where the modulation frequency was set at 5.5 Hz. The multiplicative term overlapped the 2nd harmonic.
The last graph (FIG. 7e) shows an optimised modulation frequency. Both the multiplicative term and the modulation fundamental were beyond the 3rd harmonic region. This was done to cater for the variability of the heart rate.
Due to the proximity of the multiplicative term with the harmonics and the variability of the heart rate, the optimisation would result in a fairly narrow frequency band.
As mentioned earlier above, the modulation pressure can be adjusted according to the magnitude of the multiplicative term. The modulation pressure must be high enough to cause artificial pulsations in the venous blood system and yet not too strong as to affect the arterial system. The indicator of too great a pressure is the appearance of the multiplicative term in the frequency spectrum of the detected signal. Since the multiplicative term is an indication of the degree to which the underlying arterial system is being disturbed, it is important the term is minimised.
The graph shown in FIG. 8 illustrates the effect of modulation pressure and cuff to oximeter spacing on the level of the detected signal. Each of the curves plots the measured signal level against the modulation pressure, and the black dot indicates the point at which the multiplicative term becomes significant. By moving the modulation site closer to the site of measurement, the multiplicative term can be kept small, and the detected signal of a measurable level even with a relatively low pressure modulation. Typically the modulation pressure is set to result in approximately 0.1% variation in the DC level of the detected signal. The ideal balance between modulation pressure and modulation/measurement site spacing is represented by the space denoted "ideal design window". Typically a spacing of 30 to 50 mm is appropriate, depending on the pressure applied via the cuff. If the spacing is too small, the pulsations from the transducer will be coupled to the oximeter probe and cause a motion artefact in the received signal.
It is to be appreciated that in addition to the cuff and air pump arrangement, other ways of modulating pressure to generate a venous pulsatile signal are envisaged. This could for example be achieved by direct mechanical means or by applying electrical or thermal impulses to the site of modulation. Further, whilst the embodiment described applies positive pressure to the subject, a similar effect can be achieved by the application of negative pressure, for example by providing a vacuum pump to generate the perturbations to the system.
Whilst the oximeter probe and pressure transducer will generally be separate devices, they may also be formed as an integral device, provided that mechanical coupling between the transducer and the probe is avoided.
The specific embodiment described has employed the pulse oximeter probe in a transmission mode (in other words, the light passing through the digit, light source and sensor lying on opposite sides of the digit) but a probe operating in reflection mode could also be used. Since the veins lie close to the surface of the skin, if a reflection mode probe is used, the position of the probe could be used at a wide range of locations on the body, and is not limited to those regions where light is transmissible through the body tissue (digits, ear lobes etc).
In the event that both arterial and venous oxygen saturation are measured, this can be achieved simultaneously as described above. Alternatively, should the external pressure modulations disturb the calibration of the arterial oxygenation measurement, the measurement of these two quantities can be separated either physically (by measuring arterial and venous saturation at different sites) or temporally (e.g. by multiplexing the arterial and venous measurements over time). Where the measurement sites are separated physically, the probe may be either integrally formed, e.g. incorporating an arterial oximeter, and a venous oximeter and cuff, arranged to measure from two adjacent fingers, or may be formed of separate arterial and venous monitor devices.
A non-invasive method of determining venous oxygen content has been described which will allow accurate real time monitoring at lower risk to patients. This is of particular relevance during surgical recovery and management of therapy. A number of areas where the invention will find application are outlined below.
When patients are suffering from severe illness of either the cardiovascular system or the lungs their survival depends on the ability to optimise the delivery of oxygen to their tissues. Tailoring of oxygen delivery to match a patient's requirements is very difficult, even in an intensive care unit. It relies upon a combination of clinical assessment, laboratory blood tests, haemodynamic data and oximetry. These data are obtained from the insertion of catheters into the radial/femoral arteries and pulmonary artery and can take 1-2 hours to complete. The insertion of these lines carries a significant morbidity and mortality. The integration of the data needed to tailor the patient's oxygen delivery can take several hours to achieve. As such it is not suited to rapidly changing physiological situations. A rapidly applicable, non-invasive, measure of tissue oxygen delivery would benefit all critically ill patients in intensive care units. It would also be useful for patients in High Dependency Units and ordinary hospital wards. Perhaps its most useful potential application, however, would be in the resuscitation of patients. Non-invasive venous oximetry would allow resuscitation to be more focused; since the survival of patients suffering out-of-hospital cardiac arrest is <5%, this would be a great breakthrough. In addition, non-invasive venous oximetry would prove an invaluable aid to the safe transfer of critically ill patients between intensive care units. Indeed, studies performed by the inventors have shown that there is a close correlation between SvO2 and cardiac output (CO), and therefore the method can also serve as a non-invasive indicator of Cardiac Output based on measurement of SvO2.
Cardio-Pulmonary Bypass (CPB) and Extracorporeal Membrance Oxygenation (ECMO) involve the temporary replacement of cardio-vascular and lung function by use of a pump-oxygenator. CPB is used in the operating theatre, whilst ECMO is used in the Intensive Care unit to support patients who are suffering from potentially reversible heart/lungs failure. Blood flow from the pump is non-pulsatile, and this means that conventional oxime try is ineffective. The ability to track oxygen delivery non-invasively with a device which does not rely on naturally pulsatile flow would have wide application. Since Coronary Artery Bypass Grafting (CABG) is the most commonly performed operation in the USA today this represents a significant potential application.
The diagnosis and monitoring of vascular disease and circulatory function will be enhanced by the availability of SvO2 measurement. Elevated SvO2 measures at rest may indicate reduced tissue perfusion due to impaired blood flow. Monitoring venous oxygen saturation will therefore allow the severity of injury and functional compromise resulting from trauma to be examined more easily. Depressed SvO2 measures may suggest tissue dysfunction and monitoring will allow judgements to be made about the viability of tissue during trauma or disease.
As the PaO2 drops below about 40 Torr, even small changes in the partial pressure of inspired oxygen result in large decreases in SaO2. SvO2 measurement will enhance the monitoring of hypoxia in a range of conditions. Physical training evokes peripheral adaptations to allow for effective utilisation of the increase in O2 delivery resulting in higher O2 diffusional conductance in muscle. The effect of endurance training can be observed as a greater capillary density allowing for a longer transit time at a given blood flow and also a higher a vO2 difference. The invention would be a valuable tool in performing exercise/stress tests and examining the effects of heat and cold, micro-gravity and dehydration. It would also support numerous other medical and physiological research themes.

Loughborough University Enterprises Ltd.


A method for making expandable foam patterns for use in an evaporative casting process includes assembling a plurality of mold parts of a molding tool having mold walls defining a mold cavity, each mold part having a housing wall separated from the mold wall to form a heat exchange cavity adjacent the mold wall, inserting at least one heat pipe adjacent a mold wall portion, filling the mold cavity formed by the mold parts with expandable foam beads, heating the tool by introducing a fluid heat exchange medium to the heat exchange cavity to heat the cavity for a predetermined time, cooling the tool and separating the mold parts to release the expanded foam pattern.

Patterns used in evaporative castings must be precisely made since they define the shape of the part to be made within a sand casting cavity formed by packing the pattern in sand. Such patterns are consumed during the evaporative casting process, and a large number of patterns are required for evaporative casting since each part requires a new pattern. As a result, economies and improvements in the production of the patterns can have a substantial effect upon the cost and quality of the end product produced by the evaporative casting process.
A previously known method for making such patterns comprises assembling a plurality of previously configured mold parts including mold wall surfaces that define the pattern forming chamber. The mold walls are separated apart from housing walls of each mold part to form a heat exchange cavity therebetween. The mold parts also include couplings for connecting the heat exchange cavities with fluid heat exchange such as cold water and hot steam. Assembly of the mold parts forms a pattern making tool.
To prepare expanded foam patterns with such parts, beads are formed from polystyrene or other materials such as polymethyl methacrylate (PMMA) or polyalkylene carbonate (PCA) and a thermal expanding agent such as pentane or the like. The beads, originally about the size of salt grains, are heated in a large container to a temperature of about 215.degree. F. to pre-expand the grains and form beads of a density of about 0.9-1.9 pounds per cubic foot. The pre-expanded pellets are then transferred to fill the mold cavity defined by the mold walls. The pellets are then heated for a limited time depending upon the degree of fusion desired and the composition of the pellets, for example, generally from a fraction of a minute to several minutes, by steam introduced to the heat exchange cavities at a temperature of about 215.degree. F. This heating step will expand, soften, distort and merge the expanded beads together to form a unitary foam pattern shaped by the mold walls. The expansion process is then stopped by running cold water through the heat exchange cavities. After separating the mold parts, the pattern is removed from the mold.
In such a process, temperature variations along the mold wall can have a profound effect upon the quality of the part produced. The degree of expansion of the pellets depends upon the temperature and the time duration at which the temperature is maintained. Accordingly, dimensional stability of the pattern and the surface quality of the pattern are improved when the temperature of the mold cavity surface is consistent throughout the cavity. However, when the mold wall includes recessed wall portions or various wall thicknesses, temperature along the mold surface may be inconsistent throughout the stages of operating the mold. When parts of the mold wall surface are recessed away from the heat exchange chamber, the resistance to heat energy transfer from the heat exchange medium can interfere with consistent heat exchange to the foam beads forming the pattern. Moreover, while thin mold wall portions may be desired for efficient heat transfer, a thicker mold wall portion will absorb more heat to reach the same temperature even though the heat exchange medium contacting it has a substantially constant temperature. Furthermore, a greater amount of heat energy may need to be transferred to a mold wall part which defines a mold wall cavity portion for a relatively thick portion of the pattern to be formed.
The problem of inconsistent mold wall temperatures is particularly acute in molding tools using mold cores since the cores typically do not include heat exchange passages. Accordingly, previous expanded foam patterns may require assembly of two or more expanded foam parts by gluing or the like to complete a single pattern. Such assembly interferes with the integrity of the pattern and the quality of the pattern surfaces.
Moreover, the quality of the pattern formed has a substantial effect upon the part to be cast using the pattern in an evaporative casting process. Improved surface finish and dimensional stability in the pattern can substantially reduce if not eliminate certain machining operations otherwise necessary to satisfy the molded part's required specifications. Each such machining operation which can be eliminated represents a substantial savings in energy, labor and tooling costs which must be incurred in order to produce a finished part after evaporative casting.
In order to overcome the problem of temperature variation, a previously known attempt for improving the pattern casting method included venting the core. In a hollow vented core or mold, vent passageways are provided in the mold wall forming the mold cavity so that the heat exchange chamber is coupled for fluid communication with the mold cavity. The vent typically includes restricted slots to prevent expansion of the foam material into the vent or the heat exchange chamber during molding. Before the foam material is introduced into the cavity, the vent passages permit the heat exchange fluid medium to enter the cavity and contact both sides of the mold wall defining the mold cavity to enhance uniformity of the heat transfer to the molded material. Although such venting provides greater uniformity in temperature throughout the mold cavity, foam material can expand into the slots of the vent passage during the process. The bulges of foam into the vent cause the part to deviate from the part specifications. Furthermore, they cause interference between the foam pattern and the core or mold wall surfaces and cause scoring of the pattern during removal of the core or the pattern from the mold. Such interference can cause substantial damage to the surface finish of the molded foam part and thus, to the cast workpiece formed with the pattern. As a result, a large number of unacceptable patterns are produced by such a method.
Another process for making foam patterns without the difficulties of vented cores or mold walls comprises the use of hollow cores or mold walls through which the heat exchange medium can penetrate into deeply recessed portions of the mold wall. Although such hollow openings permit the heat exchange medium to contact the mold walls, the restrictions to flow caused by dimensional limitations of such channels can cause a substantial variation in temperature between the recessed mold wall and contiguous mold wall portions. Accordingly, the heat exchange medium must remain in the heat exchange cavities for longer periods of time in order to permit the temperature at the bottom of the recess to reach the temperature of the remaining mold wall portions. Such time delays in the pattern forming process can substantially reduce the production output of the foam patterns and thus the output of the evaporative casting process. As a result, the number of tools to produce the patterns must be increased in order to provide a satisfactory production rate. Furthermore, the variation in expansion rates throughout the mold cavity affects the surface consistency and dimensional stability of the pattern.
The inability to uniformly transfer heat has also been recognized in the art of injection molding, where the molten material, usually plastic, is heated before introduction into the mold chamber. In such a process, the material must be heated to a molten state in order to be transferred from its source through delivery conduits and within the mold cavity so that it remains capable of conforming to the mold cavity shape. In order to maintain enough heat in the fluid during transfer between the source and the cavity, U.S. Pat. No. 4,034,952 to Stewart discloses a hot plastic injection bushing surrounded by a heat pipe to maintain all parts of the tubular bore at the same temperature as an externally controlled electrical heater.
U.S. Pat. Nos. 4,338,068 and 4,387,762 disclose a variable volume heat pipe for use in injection molding. U.S. Pat. No. 4,500,279 discloses a hot manifold system employing heat pipes adjacent runner channels to provide continuous heating of the runners during injection molding. Such injection molding apparatus require heat pipes to transfer sufficient heat energy to maintain the high temperatures, often 400.degree.-500.degree. F., necessary to keep the mold material fluid so that a fixed amount of molding material can be transferred in its molten state to completely fill the mold cavity.
On the other hand, foam pattern molds do not require the high temperatures required for injection molding. Moreover, the foam material is transferred without reaching a molten state. Rather, the expandable foam pellets are heated to increase the volume of the pellets within the confined space of the mold chamber. As a result, the time during which heat is supplied rather than the ability to maintain high temperatures is required in foam pattern making.
Furthermore, injection molded parts can be removed from the mold whenever the molding material solidifies. Once injection material has filled the mold cavity, the application of heat is terminated. The material solidifies once the heat energy is no longer input to maintain the molten state of the material. As a result, previously known injection molding apparatus do not address the problem of maintaining consistent temperature variation throughout all parts of the mold during the curing process as is required in the cooling phase of making expanded foam patterns.
From the foregoing, the skilled artisan will appreciate that the teachings of the previously known injection molding apparatus are not readily applicable to expanded foam pattern making processes.

Method for making evaporative casting patterns